Radiography is a long known medical diagnostic imaging technique.
In a conventional radiography system, an x-ray source is actuated to direct a divergent area beam of x-rays through a patient. A cassette containing an x-ray sensitive phosphor screen and light and x-ray sensitive film is positioned in the x-ray beam on the side of the patient opposite the source. X-radiation passing through the patent's body is thereby attenuated in various degrees to produce on the film a shadow image of a portion of the patient through which the x-rays pass.
More recently, digital radiographic techniques and systems have been developed. In digital radiography the source directs x-radiation through a patient's body to a detector assembly located in the beam path beyond the patient. The detector produces electrical signals defining the radiation pattern emergent from the patient and incident on the assembly. These signals are then processed to yield a visual display of the image.
The detector assembly includes an elongated array of individual detector elements. A detector element can suitably comprise a scintillator material positioned in front of a photodiode. Each detector element responds to x-radiation incident on the phosphor to produce an analog electrical charge signal indicative of such radiation. The set of these analog electrical signals represents the radiation pattern emergent from the patient's body.
The analog signals are sampled and processed by imaging circuitry, primarily to improve their signal to noise ratio, and are subsequently digitized.
The digital signals are fed to a digital data processing unit. The data processing unit records and/or processes and enhances the digital data.
A display unit responds to appropriate digital data representing the image to convert the digital information back into analog form and to produce a visual display of the patient's internal body structure derived from the acquired image pattern of radiation. The display unit can be coupled directly to the digital data processing unit for substantially real time imaging, or can be fed stored data from digital storage means such as tapes or disks, representing patient images acquired in earlier studies.
Digital radiography includes techniques in which a thin spread beam of x-radiation is used. In practice of this technique, often called "scan (or slit) projection radiography" (SPR), the spread beam is scanned across the patient, or the patient is movably interposed between the spread beam x-ray source and the detector assembly, the detector array being maintained in continuous alignment with the beam. The relative movement effected between the source-detector arrangement and the patient's body scans a large portion of the body.
Discrete element detectors have been proposed comprising a single line of detector elements. Other proposals have included planar rectangular detector arrays of square detector elements.
Details of certain aspects of digital radiography systems such as described here are set forth in the following publications, hereby expressely incorporated by reference:
Mattson, R. A., et al, "Design and Physical Characteristics of a Digital Chest Unit", S.P.I.E. Volume 314, Digital Radiography (1981);
Arnold, B. A. et al "Digital Radiography: An Overview" Proceedings of S.P.I.E. Volume 273, March 1981;
Kruger, R. A. et al "A Digital Video Image Processor for Real Time X-ray Subtraction Imaging" Optical Engineering, Volume 17, No. 6 (1978);
U.S. Pat. No. 4,383,327, issued on May 10, 1983, to Kruger.
European patent application Publication No. EP 0115125-A1, published Aug. 8, 1984, by Gary L. Barnes and entitled "Split Energy Level Radiation Detection";
U.S. patent application Ser. No. 542,384, filed Oct. 17, 1983 by Mattson, R. A., et al entitled "Improving Signal Characteristics in Digital Scan Projection Radiography", and owned by the assignee of this application.
U.S. patent application Ser. No. 653,955, filed by Sones, et al. on Sept. 21, 1984, entitled "Digital Radiography Detector Resolution Improvement" and owned by the assignee of this application.
U.S. patent application Ser. No. 673,779, filed on Nov. 21, 1984 and entitled "Imaging With Focused Curved Radiation Detectors" and owned by the assigneed of this application.
An important technique for enhancing a digitally represented image is called "subtraction". There are two types of subtraction techniques, one being "temporal" subtraction, the other being "energy" subtraction.
Temporal, sometimes called "mask mode" subtraction, is a technique that can be used to remove overlying and underlying structures from an image when the object of interest is enhanced by a radiopaque contrast agent. Images are acquired with and without the contrast agent present and the data representing the former image is subtracted from the data representing the latter, substantially cancelling out all but the blood vessels or anatomical regions containing the contrast agent.
A principal limitation of digital temporal subtraction is the susceptibility to misregistration, or "motion artifacts" caused by patient movement between the acquisition of the images with and without the contrast agent.
An alternative to temporal subtraction, which is less susceptible to motion artifacts, is energy subtraction. Whereas temporal subtraction depends on changes in the contrast distribution with time, energy subtraction exploits energy-related differences in attenuation properties of various types of tissues, such as the difference of the attenuation characteristics of soft tissue and bone.
Soft tissue shows less change in attenuation capability with respect to energy than does bone.
This phenomenon enables performance of energy subtraction. In practicing that technique, pulses of x-rays having alternating higher and lower energy levels are directed through the patient's body. When a lower energy pulse is so generated, the detector and associated digital processing unit cooperate to acquire and store a set of digital data representing the image produced in response to the lower energy pulse. A very short time later, when the higher energy pulse is produced, the detector and digital processing unit again similarly cooperate to acquire and store a separate set of digital information representing the image produced by the higher energy pulse. The values obtained representing the lower and higher energy images are then processed in accordance with techniques described in the following publication, hereby incorporated by reference: Lehmann, L. A. et al, "Generalized Image Combination in Dual KVP Digital Radiography" Medical Physics Volume 8, pp. 659-667 (1981). By processing in this manner, the image contrast and visibility of different tissues is substantially enhanced.
Energy subtraction has the advantage, relative to temporal subtraction, of being substantially not subject to motion artifacts resulting from the patient's movement between exposures. The time separating the lower and higher imaging acquisitions is quite short, often less than one sixtieth of a second.
An important disadvantage in dual energy subtraction techniques results from the necessity of rapidly alternating the output of an x-ray tube between high and low levels. This requirement gives rise to severe problems in a practical clinical device. The switching frequency is required to be on the order of 500 Hz. and insufficent photons (x-ray energy) result when even the highest capacity x-ray tubes are combined with realistically narrow x-ray beam slit widths and rapid scanning rates.
In order to eliminate this problem, a detector assembly has been proposed which enables the practice of energy subtraction radiography with the use of a constant output x-ray source.
In accordance with this proposal, an example of which is described in the above incorporated Barnes published European patent application, a dual layer dual energy radiation detector assembly has been suggested. A first layer comprises a rectangular planar array of square detector photodiode elements including a first radiation sensitive scintillation material overlying the photodiodes and being selected for its primary response to radiation of a lower energy range. A second planar layer is located, or "stacked", directly behind the first layer, with respect to the x-ray tube, and comprises a similar rectangular array of detector elements congruent and aligned with the first layer. The second layer includes a second radiation sensitive scintillation material selected for its propensity to respond primarily to radiation of a higher energy level, which has passed through the first layer substantially without being detected.
Such a dual energy detector structure, when used in conjunction with an x-ray tube emitting energy over a wide range, will provide data describing two separate images, i.e., one an image of lower energy x-radiation passing through the subject, the other being an image describing the pattern of higher energy radiation.
The Sones patent application referred to above describes the use of curved arrays of detector elements in digital radiography applications. One of the problems of this technology involving curved arrays relates to providing mechanical structure for holding the arrays in their desirable fixed curved geometrical relationship. In the past, independent support structure was proposed, whose only function was to provide mechanical stability.
The difficulty of maintaining the proper geometrical relationship among the detector elements becomes exacerbated in situations in which dual layer detectors are employed, because of the additional complexity of these detector arrangements. It was required to provide support structure which was both compatible with the necessary geometrical relationships to be maintained among the detector elements, and did not interfere with system operation, and which provided sufficient sturdiness to maintain detector element alignment in normal use. This situation was further complicated by the need for moving the entire detector assembly in synchronism with the thin spread beam of x-rays.
It has been proposed in computed tomography applications to employ scintillation detector elements electrically, but not mechanically, connected to printed circuitry on boards, which circuitry performs part of the image processing of signals produced by the detector elements.
The usual proposal was to couple the scintillator elements to the printed circuitry by soldered or welded terminals, or by plugs. Problems exist with connection by solder or welding. Each terminal coupling must be made individually, often by hand, a process which is tedious and costly.
Once made, such couplings could not be disassembled without destroying the coupling means, i.e., the soldered or the welded connections. This means that assemblies including such connections are difficult to repair, in that, after the repair, the couplings must be remade in their entirety.
Additionally, the brittle nature of the couplings thus made introduces durability problems wherever movement or vibration is present.
The use of plugs as connectors, while more easily disassembled than welded or soldered connections, results in the plugs taking up significant portions of the circuit board surface, which limits the placement of board circuitry, sometimes increasing the difficulty of producing the appropriate circuit boards.
Another type of problem arises from the use of radiation detector assemblies incorporating filters and phosphor materials, such as described in the Barnes published European application, is that the phosphor material and filters are permanently fastened in place when the detector assembly is manufactured. Such assemblies have no facility for simply changing phosphor or filter materials, and thus cannot be easily "tuned" for different types and levels of radiation, and for different types of studies.